RESEARCH ARTICLE Tendon material properties vary and are



RESEARCH ARTICLE Tendon material properties vary and are
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The Journal of Experimental Biology 215, 000-000
© 2012. Published by The Company of Biologists Ltd
Tendon material properties vary and are interdependent among turkey hindlimb
Andrew Matson, Nicolai Konow, Samuel Miller, Pernille P. Konow and Thomas J. Roberts*
Department of Ecology and Evolutionary Biology, Brown University, Providence, RI 02912, USA
*Author for correspondence ([email protected])
The material properties of a tendon affect its ability to store and return elastic energy, resist damage, provide mechanical
feedback and amplify or attenuate muscle power. While the structural properties of a tendon are known to respond to a variety of
stimuli, the extent to which material properties vary among individual muscles remains unclear. We studied the tendons of six
different muscles in the hindlimb of Eastern wild turkeys to determine whether there was variation in elastic modulus, ultimate
tensile strength and resilience. A hydraulic testing machine was used to measure tendon force during quasi-static lengthening,
and a stress–strain curve was constructed. There was substantial variation in tendon material properties among different
muscles. Average elastic modulus differed significantly between some tendons, and values for the six different tendons varied
nearly twofold, from 829±140 to 1479±106MPa. Tendons were stretched to failure, and the stress at failure, or ultimate tensile
stress, was taken as a lower-limit estimate of tendon strength. Breaking tests for four of the tendons revealed significant variation
in ultimate tensile stress, ranging from 66.83±14.34 to 112.37±9.39MPa. Resilience, or the fraction of energy returned in cyclic
length changes was generally high, and one of the four tendons tested was significantly different in resilience than the other
tendons (range: 90.65±0.83 to 94.02±0.71%). An analysis of correlation between material properties revealed a positive relationship
between ultimate tensile strength and elastic modulus (r20.79). Specifically, stiffer tendons were stronger, and we suggest that
this correlation results from a constrained value of breaking strain, which did not vary significantly among tendons. This finding
suggests an interdependence of material properties that may have a structural basis and may explain some adaptive responses
observed in studies of tendon plasticity.
Key words: tendon, tissue material properties, elasticity, tensile strength.
Received 19 March 2012; Accepted 28 June 2012
Tendons transmit force between muscles and the skeleton, but they
also play an important role as springs (Cavagna et al., 1977;
Alexander, 2003). Elastic deformation of tendons allows them to
store potential energy, provide mechanical feedback and amplify
or attenuate muscle power (Roberts, 2002; Konow et al., 2011;
Roberts and Azizi, 2011). In walking and running, the stretch and
recoil of tendons improves locomotor economy by reducing muscle
work and allowing muscles to operate at lengths and velocities that
are favorable for force production (Heglund et al., 1982; Alexander,
1991; Roberts et al., 1997; Lichtwark and Wilson, 2006; Sawicki
et al., 2009).
The mechanical properties of any given tendon will determine
its tendency to stretch, shorten and potentially fail under muscular
loads. These mechanical properties in turn affect muscle length
changes, joint motions and the cycling of elastic strain energy.
Presumably, some matching of tendon properties to muscle
properties allows for both effective spring-like activity as well as a
degree of protection against tendon damage or failure (Ker et al.,
1988). Variation in tendon dimensions is an important means of
tuning mechanical properties. For instance, if material properties
are held constant, a short tendon with a large cross-sectional area
will tend to be stiffer than a long thin tendon.
It is less clear to what extent the material properties of tendons
vary in response to functional demands. It is commonly stated that
all vertebrate tendons are relatively similar in elastic modulus (EM)
(Zajac, 1989; Alexander, 2003), and some comparative studies have
supported this assumption (Bennett et al., 1986; Pollock and
Shadwick, 1994). However, reported values vary widely, from low
EM values of 160MPa in juvenile pig tendon (Shadwick, 1990), to
values exceeding 2000MPa for human tendon (Maganaris et al.,
2004). It is difficult to interpret the variation in material properties
from literature values because methodological challenges associated
with these measurements leave some uncertainty as to whether
variation in reported values might be explained, in part, by
differences in measurement techniques. One particular challenge
for in vitro measurements is the technique for clamping soft tissue,
which can lead to errors in strain measurements due to slipping, as
well as underestimates of tendon strength due to damage at the
tendon–clamp interface. Recent studies have avoided problems
associated with clamp slipping by using surface markers in an intact
in situ preparation, and have demonstrated variation in EM both
between tendons of different muscles (Cui et al., 2009) and along
the length of an individual tendon (Arruda et al., 2006; Wood et
al., 2011).
We used wild turkey hindlimb muscles to determine whether
tendon material properties vary among muscles. Turkeys are a good
model because their distal hindlimbs include numerous muscles with
long free tendons. Most importantly, many of their tendons are
continuous with ossified regions that provide a means for secure
The Journal of Experimental Biology 215 (?)
clamping, avoiding many of the problems associated with clamping
soft tissues. By isolating six different muscle–tendon units (MTUs)
from the same limbs, we were able to achieve a broad sample for
comparisons of tendon EM, strength and resilience. Our primary
goal was to determine whether these properties varied among
tendons. A secondary goal was to determine whether differences in
one property correlated with differences in another.
For tendon testing, data were collected from hindlimb tendons of
captive-bred wild turkeys (Meleagris gallopavo Linnaeus 1758;
N14; mass: 3.1–4.0kg). Six different hindlimb muscles were
identified following the study of Harvey and co-workers (Harvey
et al., 1968). The tendons studied belonged to the following
muscles: gastrocnemius complex (GA; N8), tibialis cranialis (TC;
N9), extensor digitorum longus (EDL; N12), flexor perforans
digiti 3 (FPD3; N12), flexor perforans et perforatus digiti 3 (FPPD3;
N12) and flexor hallucis longus (FHL; N12) [nomenclature
follows Vanden Berge and Zweers (Vanden Berge and Zweers,
1993)]. Tendon samples were harvested from frozen specimens used
in previous studies. Samples were excluded if they were damaged
during dissection or preparation. Dimensional data for the associated
muscles were collected from eight separate individuals of
comparable weight (3.1–3.6kg). All animal use was approved by
the Brown University Institutional Animal Care and Use Committee.
Sample preparation
All tendon samples were isolated from intact hindlimbs stored at
–18°C prior to testing. Before dissection, limbs were thawed at room
temperature in zip-seal bags to avoid desiccation. During dissection,
each MTU was isolated from the origin of the muscle to the distal
insertion of the tendon using blunt dissection. The free tendons,
including both bony and soft regions, were separated from their
muscle bellies and placed in saline at room temperature. When the
dissection was complete, tendons were wrapped in saline-saturated
gauze and stored at –18°C, and thawed again on the day of
mechanical testing. One to two repeated freezing and thawing cycles,
as in the case of our study, have limited effect on tendon EM (Ker,
1981; Chen et al., 2011), whereas multiple freeze–thaw cycles may
damage tendon microstructure (Chen et al., 2011).
Tendons were isolated from hindlimbs of animals that had been
used previously in locomotor studies. There was some variation
among individuals in the extent of treadmill training they had
experienced, and the details of the locomotor regime. Some animals
were trained for uphill running, others for downhill, and others
received minimal training. These differences were controlled for
statistically as described below.
Muscle dissection and measurement were performed before tendon
testing on separate animals. Each muscle was weighed and then
longitudinally bisected to facilitate measurements of fascicle length
and pennation angle (Roberts et al., 1998). Average muscle
physiological cross-sectional area (Am) was calculated as:
Am  (Mmcos)/(mLm),
where Mm is the mass (g) of the muscle,  is the average measured
pennation angle (deg) of the muscle fibres, m is the density of
muscle, assumed to be 1.06gcm–3 (Mendez and Keys, 1960), and
Lm is the average measured fascicle length (cm).
The digital flexors and extensors all included a region of soft
tendon bounded at both the proximal and distal ends by segments
of bony tendon. For these tendons resting length (L0) was defined
as the distance from the proximal bony–soft tendon junction to the
distal bony–soft tendon junction (Fig.1). L0 was measured before
mechanical measurements while the tendon was clamped at a passive
tension of 0.8N. For the TC and GA, L0 was determined from the
distance between the marked segment of interest as determined from
video. The mass of each soft tendon was obtained after mechanical
measurements. Average tendon cross-sectional area (At) was
calculated from tendon volume and length:
At  Mt/(tL0),
where Mt is tendon mass in grams and t is the density of tendon
[1.12gcm–3 (Ker, 1981)].
Mechanical measurements
We used a hydraulic testing machine (858 Mini-bionix II, MTS
Systems, Eden Prairie, MN, USA) operating under length control
to measure tendon stiffness, resilience and breaking strength. For
the four digital flexor and extensor tendons, the calcified regions
of the tendon provided a very favorable site for clamping, as it was
possible to clamp bony tendon very firmly without significantly
damaging the tissue. Tendons were trimmed to leave approximately
1cm of calcified tendon, which was clean of any overlying soft
tissue and which allowed for clamping while preserving the
bony–soft tendon junction (Fig.1). Digital flexor tendons in the
turkey hindlimb all follow the same pattern of bony and soft regions.
Tendons are calcified at the interface with the muscle, then soft as
they cross the intertarsal (ankle) joint, followed by a bony region
that runs the length of the tarsometatarsus. Distal to the
tarsometarsus, the tendons become soft again as they cross joints
of the digits, and many interconnect via vinculae. In mature tendons,
the transition from bony to soft is an abrupt one; there does not
appear to be a significant region of transitional tissue. We used the
proximal bony region, at the muscle, and the bony region within
the tarsometatarsus to clamp our digital tendon samples, thus
measuring the soft tendon region that spanned the intertarsal joint.
The proximal end of the tendon was clamped to the top clamp, the
controllable, mobile component of the MTS machine. The distal
bony region was clamped to a stationary clamp attached to the MTS
force transducer. Sandpaper was used to increase clamp friction in
both clamps, and the surface of the bony tendon was scraped with
a razor blade to ensure that there was no overlying soft tissue.
Tendons were irrigated with physiological saline and kept wrapped
in a saline-saturated cloth in between data collections to avoid
The GA and TC tendons required a different clamping technique.
For GA, the proximal end divides into two as the bony tendons
follow the myotendinous junctions of the lateral and medial
gastrocnemius bellies. Both divisions were clamped in the upper,
mobile clamp. The distal ends of both GA and TC are soft and insert
directly into the tarsometatarsal bone. Dissection of these tendons
involved excision of approximately 1cm3 blocks of bone at the distal
insertions of these tendons onto the tarsometatarsus. The distal bony
block containing the tendon insertion was potted in a liquid plastic
compound (Smooth-Cast 300Q, Smooth-On, Easton, PA, USA) and
this complex was clamped to the bottom, stationary clamp. L0 was
defined from the proximal bony–soft junction to the point of tendon
insertion into the bone.
Force and length signals were sampled at a frequency of 1024Hz
via MTS Station Manager software. The same hardware and
software system was used to control position of the top clamp within
the MTS frame. Prior to measurements, clamp position was adjusted
Variability in tendon material properties
Fig.1. Photograph of a sample tendon from the hindlimb of the wild turkey
Meleagris gallopavo. Arrowheads indicate the bony–soft tendon junction.
Scale bar, 1cm.
Strain ( L/L0)
Data analysis
Data collected by MTS Station Manager software was imported into
IGOR Pro 6.0 (Wavemetrics, Lake Oswego, OR, USA), and a
force–length curve was constructed for each tendon. Strain () was
Stress (MPa)
Elastic modulus (MPa)
to achieve a resting tension of 0.8N, a force that corresponded to
no visible slack in the tendon. Each measurement trial consisted of
a series of 200 cycles of sinusoidal stretch–shortening excursions
at 4Hz, followed by a linear, isovelocity ramp lengthening. A
frequency of 4Hz was chosen as it is close to the loading frequency
in vivo during fast running. The sine wave amplitude was set to
yield a 4% L0 excursion for each tendon. The sine wave cycles
preconditioned the tendon to avoid changes in EM that are known
to occur with initial load cycles in isolated tendons (Schatzmann et
al., 1998). Over the last 10 cycles, resilience was calculated as the
average proportion of energy conserved between each
lengthening–shortening cycle. The ramp lengthening following the
sine wave cycles was driven at a rate of 1mms–1 for all specimens.
Because tendon lengths varied, this loading rate did not result in
identical strain rates for individual tendons, but this slight variation
was highly unlikely to affect our measurements as tendon modulus
is insensitive to loading rate across a broad range of values (Ker,
1981). Measurements taken during the ramp lengthenings were used
to calculate strain (Fig.2A), stress (Fig.2B) and continuous EM
The stress at the point of complete failure of the tendon, as
indicated by a precipitous drop to zero force (Fig.2B) and visible
tendon rupture, was taken as the ultimate tensile strength (UTS). It
should be noted that this estimate may be lower than the failure
strength in vivo, as it is impossible to ensure that force is evenly
distributed across all tendon fibrils in a clamped tendon (Ker, 2007).
Because of the potential for end-effects to distort mechanical
data from clamping of short, thick tendons (Legerlotz et al., 2010),
and a suspicion of slipping in GA and TC based on our
observations, MTS data were not used to determine length for
GA and TC. Instead, strain was determined from surface markers
tracked by high-speed video. Points used for digitizing were either
small surface markers (Reptile Sand, Zoo Med, San Luis Ospina,
CA, USA) or pins (000 Insect Pins, Austerlitz, Czech Republic).
Surface markers were glued onto the soft tendon at 5mm intervals
along the length of the tendon, and pins were inserted through
the tendon transversely in a similar distribution. Marker positions
were digitized using a DLT Data Viewer (Hedrick, 2008), then
smoothed for high-frequency noise using a quintic spline
interpolation (s.d.0.01–0.1). Digitized length changes between
the most proximal and distal of the soft tendon markers were used
to calculate strain.
One set of high-speed video measurements was also made for
two digital tendons to verify that MTS measurements accurately
tracked tendon strain. Strain measured from video was within 3%
of MTS measurements, confirming that end-effects and slipping
were minimal for these tendons. Therefore, MTS length
measurements were used to measure length and to calculate strain
for the digital tendons.
Time (s)
Fig.2. Sample strain (A) and stress (B) measurements for a turkey FPPD3
tendon. Data are for a controlled isovelocity lengthening at 1mms–1.
Continuous elastic modulus (EM; C) was measured as the derivative of
stress over strain. Maximum EM was determined by taking the peak of the
continuous EM curve (C). Breaking stress and strain were taken at the
point of maximum stress, as indicated by the dotted line.
defined as DL/L0 using MTS data for the digital tendons and digitized
data for GA and TC. Stress () was calculated as:
  F/At,
where  is stress (MPa), F is force (N) and At is the average crosssectional area of the tendon (mm2).
EM was measured as the instantaneous rate of change in stress
over change in strain:
EM  D/D.
To ensure that maximum EM was measured for each tendon, the
derivative of the stress–strain curve was graphed and the maximum
was identified. All values are presented as means ± 1 s.d. To test
the hypothesis that the tendons from different muscles differ with
respect to their material properties, we ran a multi-factorial ANOVA
with EM, resilience and UTS as dependent variables, and tendon
ID as an independent factor, and we used a mixed-model design to
control for individual and training effects. Post hoc testing was used
to establish if there were differences between the EM of different
tendons. Post hoc Tukey tests were used to reduce the risk of type
1 errors in multiple comparisons.
All tendons showed a J-shaped relationship between stress and strain
(Fig.3A), but the stress–strain relationship differed among tendons
of different muscles. Fig.3A shows the variability in stress–strain
curves for six different tendons from a single animal. For all tendons,
continuous EM reached a plateau over a region of approximately
3–6% strain (Fig.3B). This plateau region indicates the linear portion
The Journal of Experimental Biology 215 (?)
Elastic modulus (MPa)
Stress (MPa)
Elastic modulus (MPa)
0.06 0.08
Fig.3. Stress–strain relationships (A) and continuous elastic moduli (B) for
six different tendons in a single turkey hindlimb specimen. Colors in B are
as labelled in (A).
of the stress–strain curve, which is bound by the toe region at lower
strains and a region of yield at higher strains.
Maximum EM varied among tendons (Fig.4, Table1). FHL had
the highest average EM value (1479±106MPa) and GA had the
lowest (829±140MPa). Statistical comparisons between average
maximum EM values showed differences when comparing both
FHL and FPPD3 with EDL, FPD3 or GA (Tukey-corrected
ANOVA; all P<0.05). All other comparisons between averages of
maximum EM values were not significant.
Resilience was measured only for the four digital tendons where
bony tendon segments at both ends allowed for secure clamping.
Resilience measurements for these tendons showed a range of
90.65–94.02% (Table1). The only digital extensor tested, EDL, had
the lowest resilience value (90.65±0.83%), which was significantly
different from the three digital flexor tendons, FPD3, FPPD3 and
FHL (Tukey-corrected ANOVA, all P<0.05). All digital flexor
tendons had similar resilience values.
Measurements of ultimate tensile strength (UTS) were also
obtained for the four digital tendons (Fig.5, Table1). Due to
challenges associated with clamp slipping and damage at high forces,
not all GA and TC specimens were successfully stressed to failure,
and these tendons were thus omitted from our comparative analysis
of strength. FHL and FPPD3 were stronger than both EDL
(ANOVA, P<0.05 and P<0.01, respectively) and FPD3 (ANOVA,
Fig.4. Mean EM values for all tendons measured. Error bars show
standard deviation (N≥8). Values marked ʻaʼ differ from those marked ʻbʼ
(Tukey-corrected ANOVA, all P<0.05). TC (marked ʻa,bʼ) does not differ
from any other averages (Tukey-corrected ANOVA, P>0.05).
P<0.05 and P<0.01, respectively). To determine if variation in
tendon strength was related to tendon stresses in vivo, we used
muscle and tendon dimensions to calculate maximum stress-in-life,
assuming peak muscle stress of 300kPa. All UTS measurements
were higher than calculated maximum stress-in-life values (Fig.5).
A clear relationship that emerged was a positive correlation
between UTS and EM (Fig.6A). On average, more compliant
tendons failed at lower stresses. This correlation was present
whether analyzed by individual tendon trial (N45, r20.79,
P<0.0001) or for global averages for each tendon (N4, r20.92).
Individual measurements for EDL and FPD3 clustered at low EM
and low UTS, while individual FHL and FPPD3 measurements
clustered at high EM and high UTS.
By contrast, there was no significant relationship between strain
at failure and EM (Fig.6B). Digital tendons generally failed at a
strain of approximately 10% L0, and there was no correlation
between failure strain and EM for individuals (N45, r20.04,
P0.215), or tendon averages (N4, r20.17). Additionally, no
differences were found between any of the four average failure strain
We found significant variation among different tendons in EM,
resilience and UTS. These results add to an increasing body of
Table1. Mean values for mechanical and dimensional measurements
Elastic modulus
Ultimate tensil
strength (MPa)
strain (% L0)
mass (g)
length (mm)
area (mm2)
Measurements are means ± s.d.
Ultimate tensile strength (MPa)
Ultimate tensile strength (MPa)
Variability in tendon material properties
Fig.5. Average ultimate tensile strength values of four digital tendons
(N≥11). Averages marked ʻaʼ differ significantly from averages marked ʻbʼ
(Tukey-corrected ANOVA, all P<0.05). Thin grey lines within the bars
indicate the estimated maximum stress-in-life, calculated from muscle
dimensions and an assumed maximum muscle stress of 300kPa.
evidence indicating that tendon is a dynamic tissue with the potential
to modulate material properties in response to mechanical demand.
Our work also lends support to the hypothesis that some material
properties are interdependent. Specifically, we find a correlation
between EM and UTS, such that stiffer tendons are also stronger.
Variation in tendon material properties
Literature EM values for vertebrate tendon vary by an order of
magnitude, from less than 200MPa (Shadwick, 1990) to greater than
2000MPa (Maganaris et al., 2004), with a wide range of values in
between (Bennett et al., 1986; Bennett and Stafford, 1988; Shadwick,
1990; Lieber et al., 1991; Pollock and Shadwick, 1994; Maganaris
and Paul, 2002; Cui et al., 2009). Some of this variation is probably
explained by real variation in tendon properties. For example, the
very low value of 160MPa reported for juvenile pig tendons is
consistent with the observation that developing tendons appear to
be more compliant than those of adults (Shadwick, 1990). However,
it is difficult to draw conclusions about variation in EM from
literature measurements due to variation in measurement methods
and materials. Tendon tensile tests are methodologically simple but
may involve several sources of variation, notably including clamp
slipping, but also in the methods used to determine cross-sectional
area and prepare samples (Woo, 1982; Svendsen and Thomson,
1984; Legerlotz et al., 2010). Studies that use identical techniques
for several different tendons can provide reliable insight into the
extent of variation in tendon material properties. Cui and co-workers
(Cui et al., 2009) used in situ measurements to demonstrate variation
in EM among tendons of three cat hindimb muscles, the tibialis
cranialis (528MPa), gastrocnemius (516MPa) and extensor
digitorum longus (969MPa). Recent studies also present evidence
that EM varies within a tendon, with a longitudinal gradient of stiff
nearer the bone and more compliant nearer the muscle (Wood et
al., 2011). The variation in EM among different turkey hindlimb
muscles observed in the present study lends further support to the
idea that tendon modulus varies.
Our measured resilience values for the tendons of turkey hindlimb
muscles fell within the range of previously reported values. Studies
that evaluate in vitro tendon resilience across a broad spectrum of
mammals have found values within the range of 86–97% (Bennett
et al., 1986; Pollock and Shadwick, 1994). The observed pattern in
Strain (%L0)
Elastic modulus (MPa)
Fig.6. Relationship between EM and ultimate tensile strength (A), and
between EM and failure strain (B) for four different types of digital tendons.
For each individual tendon stretched to failure, the measured stress (A)
and strain (B) at which the tendon broke is plotted against the EM of the
tendon. In A, the fine dotted lines show quartile confidence limits and the
thick dotted lines show the prediction limits from least-squares regression.
the present study was that EDL, the lone digital extensor, was less
resilient than the three digital flexors, while none of the digital
flexors was different from each other (Table1). This trend is
consistent with that observed in mature pigs, where resilience was
lower in the digital extensors compared with the digital flexors [82.5
vs 90.8% (Shadwick, 1990)]. When measured by ultrasound in vivo,
human tendon resilience values tend to be lower than results from
in vitro studies [e.g. 78% (Kubo et al., 2002), 82% (Maganaris and
Paul, 2002) and 81% (Maganaris and Paul, 2000)]. Maganaris
(Maganaris, 2002) highlights advantages of the in vivo method in
comparison with isolated measurements such as avoidance of
potential artifact due to clamping and tendon freezing, and the ability
to use muscle contraction to load tendon. However, ultrasound
techniques present difficulties in defining the point of 0% strain and
inaccuracies attributable to the two-dimensional interpretation of
three-dimensional movement, factors that may limit the accuracy
of in vivo resilience measurements (Maganaris, 2002; Magnusson
et al., 2008).
Measured values of UTS were also variable among tendons of
different turkey hindlimb muscles and within the range of values
reported previously for mammals and birds. We found a range in
average breaking stress of 66.83–112.07MPa among the four digital
tendons, with FHL and FPPD3 breaking at significantly higher stress
The Journal of Experimental Biology 215 (?)
than FPD3 and EDL. Our values are within the range measured
previously for turkey tendons [54–117MPa (Bennett and Stafford,
1988)], mammal tendons [62–107MPa (Bennett et al., 1986) and
40–90MPa (Shadwick, 1990)] and turkey aponeurosis [53MPa
along the longitudinal axis (Azizi et al., 2009)].
It has been suggested that UTS values should be interpreted as
a lower limit, as challenges associated with clamp slippage at high
forces or tendon damage at the site of clamping can lead to
underestimates of breaking force (Bennett et al., 1986). While
clamping bony tendon segments and using video to measure tendon
strain allowed us to avoid problems associated with stress
concentrations at the clamps and slipping, our estimates of ultimate
tensile strength are still probably subject to problems of uneven
distribution of force across fibrils. The obvious yield region of the
stress–strain curves (Figs2, 3) may be explained in part by the
sequential failure of many independent fibrils. Although we made
efforts to ensure that our clamping oriented the bony and soft tendon
longitudinally, it is possible that tendon alignment in our preparations
increased the chance of uneven stress concentrations transversely
across the tendon. Thus our estimates of UTS should be taken as a
lower limit for the failure stress of tendon in vivo.
Relationship between stiffness and strength
We found a positive correlation between UTS and EM across the
range of tendons measured. By contrast, there was no significant
relationship between EM and breaking strain. There was variability
among different samples, but on average all samples broke at a strain
of close to 10% L0. We hypothesize that the correlation between
material strength and stiffness results from a relatively constrained
value of breaking strain. If a tendon can stretch by only a certain
strain before failure, then differences in EM must necessarily be
associated with differences in breaking strength. If this is the case,
then more compliant tendons will also be weaker. This positive
correlation between stiffness and strength has been observed
previously among avian tendons (Bennett, 1995). It has also been
demonstrated in an analysis that included both calcified and
uncalcified turkey tendons: calcified tendons are approximately 10
times stiffer than non-calcified tendons, and also break at higher
stresses (Bennett and Stafford, 1988).
Support for a hypothesized mechanistic link between material
stiffness and strength comes from both interventional and
comparative studies. Limb-immobilized rats show a decrease in both
material strength and material stiffness when compared with controls
(Matsumoto et al., 2003). Patterns of change in tendon material
properties with ontogeny and ageing also tend to support the idea
that EM and UTS are coupled (reviewed in Narici et al., 2008). The
tendons of young pigs are more compliant and weaker compared
with those of mature pigs (Shadwick, 1990). A study of rabbit
Achilles tendon showed that immature animals have low stiffness
(281MPa) and strength (23.9MPa) compared with mature animals
(stiffness 618MPa and strength 67.3MPa), yet both broke at similar
ultimate strains [15.7 and 16.3% for immature and mature,
respectively (Nakagawa et al., 1996)].
Advances in ultrasound methods have made it possible to
measure EM in vivo, and to track changes in modulus longitudinally
throughout an exercise intervention. This approach has in some cases
shown dramatic changes in EM, including one study that measured
a 69% increase in EM in response to exercise in an elderly human
population (Reeves et al., 2003). Breaking stress cannot of course
be estimated in vivo, but if breaking stress and material stiffness
are correlated, then the large change in EM observed by Reeves
and co-workers (Reeves et al., 2003) would also mean a similarly
large increase in breaking stress. It seems reasonable to speculate
that observed changes in EM with exercise interventions may simply
be secondary to an adaptive response to increase damage resistance.
It has also been proposed that changes in EM with training are related
to adaptive responses to reduce tendon susceptibility to fatigue
damage (Buchanan and Marsh, 2001).
Physiological basis for variation in tendon properties
The mechanisms underlying variation in tendon material properties
– both within an individual tendon in response to the mechanical
environment, or between tendons of different muscles – are not
entirely clear (Kjaer, 2004). Both the collagen fiber backbone, as
well as other components of the extracellular matrix, have been
proposed to contribute to variation in tendon material properties.
Exercise studies in animals have demonstrated that increased tendon
stiffness is associated with increased collagen density (Woo et al.,
1980) and with alterations in collagen fibril crimp angle (Wood et
al., 1988). Collagen turnover rates respond chronically to training
(Langberg et al., 2001) and acutely to exercise (Langberg et al.,
1999), along with up-regulation of other proteins and proteoglycans
(Miller et al., 2005). Other factors that may contribute to the
variability of tendon material properties include the presence of
certain proteoglycans (Robinson et al., 2005), and changes in the
chemical environment (e.g. acidity) of tendon extracellular matrix
(Grant et al., 2009). Fibrocartilage is incorporated in many tendons,
typically in regions of compressive load and at entheses (Benjamin
and Ralphs, 1998). The tendons tested in this study wrap around
the ankle joint, thus fibrocartilage elements may contribute to the
observed variation in material properties. The response of various
growth factors and signaling molecules to repetitive loading
(reviewed by Kjaer, 2004), as well as increased tendon metabolic
activity with exercise (Kalliokoski et al., 2005; Bojsen-Moller et
al., 2006), further highlight the dynamic nature of this tissue.
Our results support the idea that material properties of tendons vary
among muscles. The observed positive relationship between EM
and UTS suggest that although tendon may be plastic and adaptable
to varying mechanical environments, adaptive responses may
ultimately be constrained by coupling between different material
properties. Whether there is a structural basis for this proposed
coupling remains to be determined.
We would like to thank J. Cheney for technical assistance and A. Horner, R. L.
Marsh and members of the Evolutionary Morphology group at Brown University
for valuable feedback on early manuscript versions. We are grateful to an
anonymous reviewer whose incisive comments led to substantial revisions that
improved both the experimental design and the manuscript.
This work was supported by National Institutes of Health grant AR055295 to
T.J.R. and by the Brown Medical School Summer Assistantship Program.
Deposited in PMC for release after 12 months.
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